One of the most heavily investigated analyte sensing devices is the implantable glucose device for detecting glucose levels in patients with diabetes. Despite the increasing number of individuals diagnosed with diabetes and recent advances in the field of implantable glucose monitoring devices, currently used devices are unable to provide data safely and reliably for long periods of time (for example, months or years). See Moatti-Sirat et al., Diabetologia, 35:224-30 (1992). There are two commonly used types of implantable glucose sensing devices. These types include those that are implanted intravascularly and those that are implanted in tissue.
With reference to conventional devices that can be implanted in tissue, a disadvantage of these devices is that they tend to lose their function after the first few days to weeks following implantation. While not wishing to be bound by any particular theory, it is believed that this loss of function is due to the lack of direct contact with circulating blood to deliver sample to the tip of the probe of the implanted device. Because of these limitations, it has previously been difficult to obtain continuous and accurate glucose level measurements. However, such information is often extremely desirable to diabetic patients in ascertaining whether immediate corrective action is needed in order to adequately manage their disease.
Some medical devices, including implantable analyte measuring-devices, drug delivery devices, and cell transplantation devices require transport of solutes across the device-tissue interface for proper function. These devices generally include a membrane, herein referred to as a “cell-impermeable membrane” or “bioprotective membrane” which encases the device or a portion of the device to prevent access by host inflammatory or immune cells to sensitive regions of the device.
A disadvantage of cell-impermeable membranes is that they often stimulate a local inflammatory response, called the foreign body response (FBR) that has long been recognized as limiting the function of implanted devices that require solute transport. Previous efforts to overcome this problem have been aimed at increasing local vascularization at the device-tissue interface, but have achieved only limited success.
FIG. 1 is a schematic drawing that illustrates a classical FBR to a conventional cell-impermeable synthetic membrane 10 implanted under the skin. There are three main layers of a FBR. The innermost FBR layer 12, adjacent to the device, is composed generally of macrophages and foreign body giant cells 14 (herein referred to as the “barrier cell layer”). These cells form a monolayer of closely opposed cells over the entire surface of a microscopically smooth membrane, a macroscopically smooth (but microscopically rough) membrane, or a microporous (i.e., average pore size of less than about 1 μm) membrane. A membrane can be adhesive or non-adhesive to cells. However, its relatively smooth surface causes the downward tissue contracture 21 (discussed below) to translate directly to the cells at the device-tissue interface 26. The intermediate FBR layer 16 (herein referred to as the “fibrous zone”), lying distal to the first layer with respect to the device, is a wide zone (about 30 to 100 μm) composed primarily of fibroblasts 18, fibrous matrixes, and contractile fibrous tissue 20. The organization of the fibrous zone, and particularly the contractile fibrous tissue 20, contributes to the formation of the monolayer of closely opposed cells due to the contractile forces 21 around the surface of the foreign body (for example, membrane 10). The outermost FBR layer 22 is loose connective granular tissue containing new blood vessels 24 (herein referred to as the “vascular zone”). Over time, this FBR tissue becomes muscular in nature and contracts around the foreign body so that the foreign body remains tightly encapsulated. Accordingly, the downward forces 21 press against the tissue-device interface 26, and without any counteracting forces, aid in the formation of a barrier cell layer 14 that blocks and/or refracts the transport of analytes 23 (for example, glucose) across the tissue-device interface 26.
A consistent feature of the innermost layers 12, 16 is that they are devoid of blood vessels. This has led to widely supported speculation that poor transport of molecules across the device-tissue interface 26 is due to a lack of vascularization near the interface. See Scharp et al., World J. Surg., 8:221-229 (1984); and Colton et al., J. Biomech. Eng., 113:152-170 (1991). Previous efforts to overcome this problem have been aimed at increasing local vascularization at the device-tissue interface, but have achieved only limited success.
Although local vascularization can aid in sustenance of local tissue over time, the presence of a barrier cell layer 14 prevents the passage of molecules that cannot diffuse through the layer. For example, when applied to an implantable glucose-measuring device, both glucose and its phosphorylated form do not readily transit the cell membrane. Consequently, little glucose reaches the implant's membrane through the barrier cell layer. The known art purports to increase the local vascularization in order to increase solute availability. See Brauker et al., U.S. Pat. No. 5,741,330. However, it has been observed by the inventors that once the monolayer of cells (barrier cell layer) is established adjacent to a membrane, increasing angiogenesis is not sufficient to increase transport of molecules such as glucose and oxygen across the device-tissue interface 26. In fact, the barrier cell layer blocks and/or refracts the analytes 23 from transport across the device-tissue interface 26.
The continuous measurement of substances in biological fluids is of interest in the control and study of metabolic disorders. Electrode systems have been developed for this purpose whereby an enzyme-catalyzed reaction is monitored (e.g., by the changing concentrations of reactants or products) by an electrochemical sensor. In such electrode systems, the electrochemical sensor comprises an electrode with potentiometric or amperometric function in close contact with a thin layer containing an enzyme in dissolved or insoluble form. Generally, a semipermeable membrane separates the thin layer of the electrode containing the enzyme from the sample of biological fluid that includes the substance to be measured.
Electrode systems that include enzymes have been used to convert amperometrically inactive substances into reaction products, which are amperometrically active. For example, in the analysis of blood for glucose content, glucose (which is relatively inactive amperometrically) can be catalytically converted by the enzyme glucose oxidase in the presence of oxygen and water to gluconic acid and hydrogen peroxide. Tracking the concentration of glucose is possible since for every glucose molecule converted a proportional change in either oxygen or hydrogen peroxide sensor current will occur [U.S. Pat. Nos. 4,757,022 and 4,994,167 to Shults et al., both of which are hereby incorporated by reference. Hydrogen peroxide is anodically active and produces a current that is proportional to the concentration of hydrogen peroxide, which is directly related to the concentration of glucose in the sample. See, e.g. Updike et al., Diabetes Care, 11:801-807 (1988).
Despite recent advances in the field of implantable glucose monitoring devices, presently used devices are unable to provide data safely and reliably for long periods of time (e.g. months or years). See, e.g. Moatti-Sirat et al., Diabetologia 35:224-30 (1992). For example, Armour et al., Diabetes 39:1519-26 (1990), describes a miniaturized sensor that is placed intravascularly, thereby allowing the tip of the sensor to be in continuous contact with the blood. Unfortunately, probes that are placed directly into the vasculature put the recipient at risk for thrombophlebosis, thromboembolism, and thrombophlebitis.
Currently available glucose monitoring devices that can be implanted in tissue (e.g. subcutaneously) are also associated with several shortcomings. For example, there is no dependable flow of blood to deliver sample to the tip of the probe of the implanted device. Similarly, in order to be effective, the probe consumes some oxygen and glucose, but not enough to perturb the available glucose which it is intended to measure; subcutaneously implanted probes often reside in a relatively stagnant environment in which oxygen or glucose depletion zones around the probe tip can result in erroneously low measured glucose levels. Finally, the probe can be subject to “motion artifact” because the device is not adequately secured to the tissue, thus contributing to unreliable results. Partly because of these limitations, it has previously been difficult to obtain accurate information regarding the changes in the amounts of analytes (e.g. whether blood glucose levels are increasing or decreasing); this information is often extremely desirable, for example, in ascertaining whether immediate corrective action is needed in the treatment of diabetic patients.
There is a need for a device that accurately and continuously determines the presence and the amounts of a particular analyte, such as glucose, in biological fluids. The device should be easy to use, be capable of accurate measurement of the analyte over long periods of time, and should not readily be susceptible to motion artifact.